1. Field of the Invention
This invention relates generally to methods and apparatus for performing ultrasonic diagnosis of a target body. More particularly, the invention pertains to methods and apparatus for the measurement of sound speed in a target body. The invention is especially concerned with techniques for enhancing the accuracy of sound velocity measurements in compressible targets using one or more ultrasonic transducers in pulse-echo mode.
2. Description of Related Art
Traditional ultrasonic diagnosis is achieved by transmitting ultrasonic energy into a target body and generating an image from the resulting echo signals to survey anatomical structures. A transducer is used to both transmit the ultrasound energy and to receive the echo signals. During transmission, the transducer converts electrical energy into mechanical vibrations. Acquired echo signals produce mechanical oscillations in the transducer which are reconverted to electrical signals for amplification and recognition.
A plot or display (e.g., on an oscilloscope, etc.) of the electrical signal amplitude vs. echo arrival time yields the amplitude line (A-line) or echo sequence corresponding to a particular ultrasonic transmission. When the A-line is displayed directly as a sinusoidal pattern modulating at radio frequency (RF) it is referred to as an RF or "undetected" signal. For imaging, the A-line is often demodulated to a non-RF or "detected" signal.
Ultrasound techniques have been extensively used in the field of diagnostic medicine as a non-invasive means of analyzing the properties of tissue in vivo (i.e., living). A human or animal body represents a non-homogenous medium for the propagation of ultrasound energy. Acoustic impedance changes at boundaries of varying density and/or sound speed within a target body. A portion of the incident ultrasonic beam is reflected at these boundaries. Inhomogeneities within the tissue form lower-level scatter sites that result in additional echo signals. Images may be generated from this information by modulating the intensity of pixels on a video display in proportion to the intensity of echo sequence segments from corresponding points within the target body.
Conventional imaging techniques are widely used to evaluate various diseases within organic tissue. Imaging provides information concerning the size, shape and position of soft tissue structures using the assumption that sound velocity within the target is constant. Qualitative tissue characterization is carried out by interpretation of the grey scale appearance of the echograms. Qualitative diagnosis largely depends on the skill and experience of the examiner as well as system characteristics. However, images based only on relative tissue reflectivity cannot be used for a quantitative assessment of disease states.
Techniques for quantitative tissue characterization using ultrasound are needed for more accurate diagnosis of disease. One of the most promising parameters for quantitative measurement is sound speed. Speed of sound changes within regions of varying density and/or molecular compressibility within the tissue. Thus, it is expected that changes in tissue density due to disease will result in changes in the speed of sound. Indeed, it has been shown that changes in the speed of sound in tissue often correlate with tissue pathology. For example, cirrhotic liver tissue has been observed to contain more fat than normal liver tissue. The velocity of sound in cirrhotic tissue would therefore be expected to be lower than in normal tissue. Similarly, changes in tissue density in the region of tumors may result in changes in sound velocity in the tumor region. Unfortunately, however, such changes are relatively small and account for up to only 10% of the speed of sound in normal tissue. Therefore, accuracy in sound velocity estimation is extremely important in the analysis of tissue for pathological conditions. Usually, the accuracy of sound velocity estimations must be at least 1.0% to have specific value for quantitative tissue characterization. Hence, a need exists for the accurate measurement of sound velocity in organic tissue for clinical diagnosis.
Traditionally, measurement of sound speed has been conducted with transmission techniques. A first method of sound velocity measurement involves the transmission of sound pulses through tissue regions of known dimension and recording the time required for the pulse to traverse the region. The quotient of travel distance and travel time is computed to yield the velocity. However, due to the softness of most tissues, the dimensions of the tissue sample cannot be accurately measured which results in an error-prone measurement of sound velocity. Moreover, a reference liquid with a known speed of sound may be required to calibrate the apparatus.
A second transmission technique that has been used in medical diagnostics involves a transmitting transducer and a separate receiving transducer arranged so that they are aimed at one another with their respective axes of radiation coincident. The body of the subject is placed between the transmitting and receiving transducers. However, in vivo application of this technique has been limited to accessible organs like the breast or testes; other in vivo applications can be adversely affected by such factors as bowel gas, bone and inaccessibility.
A third transmission technique is disclosed by Ophir and Lin, "A Calibration-Free Method for Measurement of Sound Speed in Biological Tissue Samples", IEEE Transactions on Ultrasonics, FerroeIectrics, and Frequency Control, Vol. 35, No. 5, (1988) 573-577. This method allows accurate measurement of the speed of sound in soft tissue samples, while overcoming the limitations of initial techniques. The method employs a receiving hydrophone and a transmitting transducer that are coaxially aligned opposite each other. The transmitting transducer is in contact with the tissue sample, while the hydrophone penetrates the tissue sample at well-controlled incremental depths. The transit times of the pulse are recorded for all penetration depths of the hydrophone. These transit times are then plotted against the relative depths of the hydrophone, and a linear regression fit is made to the data. The slope of the fitted line is c.sup.-1, where c is the estimated speed of sound in the tissue sample. The technique requires neither calibration involving a reference medium, nor the knowledge of the thickness of the tissue sample. Yet, while this technique is capable of accurate measurements of tissue in vitro, it is clearly not suitable for speed of sound estimations in vivo.
Several techniques have been proposed for the measurement of sound velocity in vivo using ultrasonic transducers in pulse-echo mode. In one method, sound speed is measured using misregistration between pulse-echo images of the same structure obtained with two different sound beams. Sound velocity is determined from the difference in position of the same feature in different images. This method works best when a well-defined feature is available. In simulated tissue regions, known as "phantoms", thin wire added to the region will provide such a well-defined feature. However, well defined features are not easy to find in living tissue and the resulting sound speed measurement is therefore not as accurate. See Robinson et al., "Measurement of Velocity of Propagation from Ultrasonic Pulse-Echo Data", Ultrasound in Med. & Biol., Vol. 8, No. 4, (1982) 413-320.
In another pulse-echo technique called the "focus adjustment method", the mean sound speed between a reflector and linear array transducer is measured using the following three parameters: time of flight, time of flight difference, and distance between two receiver elements. To detect time of flight, the system delay-line time compensator is adjusted to obtain the sharpest reflector image. Thus, the sharpness of the target is maximized by interactive user control of signal delays at the transducer aperture. However, irregular tissue structures cause random refractions of the ultrasonic beams and make sharp focusing difficult. Also, the method is highly dependent on qualitative judgment. See Hayashi et al., "A New Method of Measuring In vivo Sound Speed in the Reflection Mode", J. Clin. Ultrasound, Vol. 16, (1988) 87-93.
A third pulse-echo method described in U.S. Pat. No. 4,669,482, involves in vivo sound velocity estimation by identifying segments of different sound velocity along a tracked ultrasonic beam using at least two widely-separated acoustic vantage points. The tracked beam is partitioned into at least two contiguous segments, the boundary between the two segments being the inner body of the body wall fat. A plurality of ultrasound pulse travel time measurements are made, each with a different apparent angle of intersection between the tracked beam and the tracking beam. For each measurement, techniques are employed for correcting refraction occurring in a transverse plane. Data pairs collected in the plurality of measurements are fitted to an appropriate equation using curve-fitting techniques well known in the art, by which the index of refraction at the body wall inner boundary, the inclination of the inner boundary, and the speed of sound in the internal tissue are derived. This technique, however, is not desirable in clinical settings because of the large "footprint" of the apparatus on the patient that results in a cumbersome examination procedure. Also, inaccuracies due to bone and/or bowel gases are common because of the wide spacing between transmitting and receiving transducers.
Hence, all the above pulse-echo techniques are clinically limited due to the need to use two widely separated acoustic vantage points and/or by the requirement that an identifiable, discrete target be available in the tissue. The use of two widely separated vantage points makes the apparatus and the examination procedure cumbersome, while the existence of a discrete target cannot always be guaranteed. Another potential problem is due to the effects of the overlying fat layer of the body on the estimation.